The field of the invention is magnetic resonance imaging (MRI) and, in particular, gradient coils for use with MRI systems.
In MRI, a uniform polarizing magnetic filed B.sub.o is applied to an imaged object along the z-axis of a Cartesian coordinate system, the origin of which is approximately centered within the imaged object. The effect of the magnetic field B.sub.o is to align the object's nuclear spins along the z-axis.
In response to a radio frequency (RF) excitation signal of the proper frequency, oriented within the x-y plane, the nuclei precess about the z-axis at their Larmor frequencies according to the following equation: EQU F=.gamma.B.sub.o ( 1)
where F is the Larmor frequency, .gamma. is the gyromagnetic ratio which is constant and a property of the particular nuclei, and B.sub.o is the polarizing field strength.
Water, because of its relative abundance in biological tissue and the properties of its nuclei, in of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 kHz/gauss and therefore, in a 1.5 Tesla polarizing magnetic field B.sub.o, the resonant or Larmor frequency is approximately 63.9 MHz.
In a typical imaging sequence, the RF excitation signal is centered at the Larmor frequency F and applied to the imaged object at the same time as a magnetic field gradient G.sub.z is applied. The gradient G.sub.z varies the strength of the magnetic field B.sub.o along the z-axis and, therefore, causes only the nuclei in a single slice through the object along an x-y plane to have the resonant frequency F and to be excited into resonance.
After the excitation of the nuclei in this slice, similar magnetic field gradients are applied along the x and y axes. The gradient along the x-axis, G.sub.x, causes the nuclei to precess at different frequencies, depending on their position along the x-axis, that is, G.sub.x spatially encodes the processing nuclei by frequency. The y-axis gradient, G.sub.y, is incremented through a series of values and encodes y position into the rate of change of phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
A weak nuclear magnetic resonance signal generated by the precessing nuclei may be sensed by the RF antenna "coil" and recorded as an NMR signal. Typically, the NMR signal is detected along two perpendicular axes to produce a quadrature signal having a real and an "imaginary" part. From this quadrature NMR signal, a slice of image may be derived according to well-known reconstruction techniques. A basic overview NMR image reconstruction is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith.
The polarizing magnetic field B.sub.o, for field strengths above approximately 0.2 Tesla, is typically produced by superconducting coils arranged along the z-axis and around a bore tube. The field is adjusted to be highly homogeneous in a spherical volume centered within the bore tube.
Gradient coils, for impressing the magnetic gradients, G.sub.x, G.sub.y and G.sub.z, on the uniform magnetic field B.sub.o, are ordinarily affixed to the bore tube. Strong repulsive forces are generated between each gradient coil and, therefore, the gradient coils are typically firmly attached to the bore tube and restrained by laminated epoxy and glass fiber. The gradient coil restraints resist such forces and reduce the acoustic noise generated by the flexing of the gradient windings.
An RF coil is also affixed to the bore tube and may be a cage like antenna having end-loops interconnected by a series of linear segments spaced circumferential about the end-loop. Such coils are taught, for example, in U.S. Pat. Nos. 4,694,255, 4,692,705 and 4,680,548, and are incorporated herein by reference.
The polarizing, gradient, and the RF coils are positioned on the outside of the bore tube so as not to interfere with placement of the patient in the bore tube for scanning. For maximum flexibility in medical applications, the bore tube is made large enough to permit the patient's entire body to be positioned within the bore tube taking into account expected variation in body dimensions between patients.
In order to practice certain pulse sequences it is desirable to greatly increase the strength and speed of the gradients G.sub.x, G.sub.y and G.sub.z. Gradient "speed" is the time required to change the magnitude of the gradient field between particular values. For most imaging techniques, higher gradient strength and response speed will decrease the time needed to acquire the NMR data required for an image. In particular, faster gradients reduce the time required to complete the MRI gradient pulse sequences and stronger gradients decrease the time needed to sample the received NMR signal by increasing the bandwidth of the NMR signal. Particularly in echo-planar imaging, where a single excitation produces a series of echoes, which are gradient encoded to generate image data, high gradient strength and speed is necessary to realize the full potential of rapid acquisition promised by this technique. Stronger gradients also increase the spatial resolution of the imaging process permitting smaller voxels to be discerned.
Also, for a number of specialized imaging techniques, higher gradient strength increases the "contrast" of the acquired data. This is true in flow and diffusion studies where the received NMR signal indicates the rate of flow of blood or other material, and in spectrographic studies, to measure the chemical shift between tissues caused by differing values of .gamma..
Both the response speed and the strength of the gradient field may be increased by increasing the power applied to the gradient coils. For a given geometry of gradient coil having a fixed inductance, the response time of the gradient coil, i.e., the amount of time required for the coil to reach a particular field strength, will depend on the applied current, and hence, on the available power. The power applied to the gradient coil is typically provided by dedicated gradient amplifiers. Therefore, increasing the maximum power that may be applied to a gradient coil is accomplished by increasing the power of each amplifier or by stacking additional amplifiers together.
There are practical limits to increasing gradient speed and strength by increasing the power applied to the gradient coil. The first limit is the cost of the gradient amplifiers. A factor of 5 to 50 increase in gradient power, over that required for conventional imaging, may be required, and the cost of the amplifiers needed to produce this power is prohibitive. The second limit is the power dissipation of the gradient coil. High gradient fields may require currents as high as 1000 amperes and such current levels create significant coil cooling problems. Cooling techniques such as circulating refrigerant among the gradient coil windings may be used, but add significantly to the cost of the MRI system.
In co-pending U.S. patent application Ser. No. 07/743,550, filed on Aug. 12, 1991 now U.S. Pat. No. 5,185,576, and entitled "Local Gradient Coil" gradient speed is increased without a corresponding increase in gradient power by decreasing the size of a gradient coil. The local gradient coil described therein provides one axis of the gradient field without undue interference with the RF field produced by an associated local and, without unduly interfering with patient access and comfort. In many pulse sequences, however, it is desirable to provide two orthogonal high speed gradient magnetic fields.